In mammals the heart is the organ responsible for maintaining an adequate supply of blood, and hence of oxygen and nutrients, to all parts of the body. Reverse flow of blood through the heart is prevented by four valves which serve as the inlet and outlet of each of the two ventricles, the pumping chambers of the heart.
Dysfunction of one or more of these valves can have serious medical consequences. Such dysfunction may result from congenital defects, or from disease induced damage. Forms of dysfunction include stenosis (reduction in the orifice of the open valve) and regurgitation (reverse flow through the closing or closed valve), either of which increases the work required by the heart to maintain the appropriate blood flows to the body.
In many cases the only effective solution is to replace the malfunctioning valve. A valve replacement operation is expensive and requires specialised facilities for open heart surgery. Replacement of failed artificial heart valves carries increased risk over the initial replacement, so there are practical limits on the number of times reoperation can be undertaken. Consequently, the design and materials of an artificial valve must provide for durability of the valve in the patient. The artificial valve must also operate without high pressure gradients or undue reverse flow during closing or when closed, because these are the very reasons for which a replacement of the natural valve is undertaken.
Mechanical valves, which use a ball or a disc or a pair of pivoting rigid leaflets as the opening member(s) can meet these combined requirements of hemodynamic performance and durability. Unfortunately, a patient who has had a mechanical valve implanted must be treated with anticoagulants, otherwise blood will clot on the valve. Clotting on the valve can either restrict the movement of the valve opening member(s), impairing valve function, or can break free from the valve and obstruct blood vessels downstream from the valve, or both. A patient receiving a mechanical valve will be treated with anticoagulants for life.
Valves excised from pigs and treated with glutaraldehyde to crosslink and stabilise the tissue are also used for replacement of defective valves. These may be mounted on a more or less rigid frame, to facilitate implantation, or they may be unmounted and sewn by the surgeon directly to the vessel walls at operation. A further type of valve replacement is constructed from natural tissue, such as pericardium, treated with glutaraldehyde and mounted on a frame. Valves from pigs or made from other animal or human tissue are collectively known as tissue valves. A major advantage of tissue valves over mechanical valves is that they are much less likely to provoke the blood to clot, and so patients receiving tissue valves are not normally given anticoagulants other than during the immediate post operative period. Unfortunately, tissue valves deteriorate over time, often as a result of calcification of the crosslinked natural tissue. This deterioration presents a problem, particularly in young patients. Thus, although the recipient of a tissue valve is not required to take anticoagulants, the durability of tissue valves is less than that of mechanical valves.
In third world countries, where rheumatic fever is still common, the problems of valve replacement in young patients are considerable. Anticoagulants, required for mechanical valves, are impractical and accelerated calcification of tissue valves precludes their use.
In the Western world, life expectancy continues to increase, and this results in a corresponding rise both in patients requiring cardiac valve replacement, and in those patients needing replacement of deteriorating artificial valves implanted in the past. There is, therefore, a need for a replacement heart valve with good hemodynamics, extended durability and having sufficiently low risk of inducing clotting so that anticoagulants are not necessary.
The natural heart valves use thin flexible tissue leaflets as the closing members. The leaflets move readily out of the orifice as blood begins to flow through the valve so that flow through the open valve is unrestricted by the leaflets. Tissue valves function similarly, providing a relatively unrestricted orifice when the valve is open. For mechanical valves, on the other hand, the closing member rotates in the orifice, but is not removed from the orifice when the valve opens. This provides some restriction to flow, but more importantly, disturbs the blood flow patterns. This disturbance to the flow is widely held to initiate, or at least to contribute significantly to, the observed tendency of mechanical valves to produce clotting.
A number of trileaflet polyurethane valve designs have been described.
A valve design, comprising a leaflet geometry which was elliptical in the radial direction and hyperbolic in the circumferential direction in the closed valve position, with leaflets dip-coated from non-biostable polyurethane solutions onto injection-molded polyurethane frames has attained durabilities in excess of 800 million cycles during in vitro fatigue testing (Mackay T G, Wheatley D J, Bernacca G M, Hindle C S, Fisher A C. New polyurethane heart valve prosthesis: design, manufacture and evaluation. Biomaterials 1996; 17:1857-1863; Mackay T G, Bernacca G M, Wheatley D J, Fisher A C, Hindle C S. In vitro function and durability assessment of a polyurethane heart valve prosthesis. Artificial Organs 1996; 20:1017-1025; Bernacca G M, Mackay T G, Wheatley D J. In vitro function and durability of a polyurethane heart valve: material considerations. J Heart Valve Dis 1996; 5:538-542; Bernacca G M, Mackay T G, Wilkinson R, Wheatley D J. Polyurethane heart valves: fatigue failure, calcification and polyurethane structure. J Biomed Mater Res 1997; 34:371-379; Bernacca G M, Mackay T G, Gulbransen M J, Donn A W, Wheatley D J. Polyurethane heart valve durability: effects of leaflet thickness. Int J Artif Organs 1997; 20:327-331). However, this valve design became unacceptably stenotic in small sizes. Thus, a redesign was effected, changing the hyperbolic angle from the free edge to the leaflet base, and replacing the injection-molded frame with a rigid, high modulus polymer frame. This redesign permitted the use of a thinner frame, thus increasing valve orifice area. This valve design, with a non-biostable polyurethane leaflet material, was implanted in a growing sheep model. Valve performance was good over the six month implant period, but the region close to the frame posts on the inflow side of the valve, at which full leaflet opening was not achieved, suffered a local accumulation of thrombus (Bernacca G M, Raco L, Mackay T G, Wheatley D J. Durability and function of a polyurethane heart valve after six months in vivo. Presented at the XII World Congress of International Society for Artificial Organs and XXVI Congress of the European Society for Artificial Organs, Edinburgh, August 1999. Wheatley D J, Raco L, Bernacca G M, Sim I, Belcher P R, Boyd J S. Polyurethane: material for the next generation of heart valve prostheses? Eur. J. Cardio-Thorac. Surg. 2000; 17; 440-448). This valve design used non-biostable polyurethane, which had tolerable mechanical durability, but which showed signs of polymer degradation after six months in vivo.
International Patent Application WO 98/32400 entitled “Heart Valve Prosthesis” discloses a similar design, i.e., closed leaflet geometry, comprising essentially a trileaflet valve with leaflets molded in a geometry derived from a sphere towards the free edge and a cone towards the base of the leaflets. The spherical surface, defined by its radius, is intended to provide a tight seal when the leaflets are under back pressure, with ready opening provided by the conical segment, defined by its half-angle, at the base of the leaflets. Were the spherical portion located at the leaflet base it is stated that this would provide an advantage in terms of the stress distribution when the valve is closed and under back pressure.
U.S. Pat. No. 5,376,113 (Jansen et al.) entitled “Closing Member Having Flexible Closing Elements, Especially a Heart Valve” issued Dec. 27, 1994 to Jansen et al. discloses a method of producing flexible heart valve leaflets using leaflets attached to a base ring with posts extending from this upon which the leaflets are mounted. The leaflets are formed with the base ring in an expanded position, being effectively of planar sheets of polymer, which become flaccid on contraction of the ring. The resulting valve is able to maintain both a stable open and a stable closed position in the absence of any pulsatile pressure, though in the neutral unloaded position the valve leaflets contain bending stresses. As a consequence of manufacturing the valve from substantially planar sheets, the included angle between the leaflets at the free edge where they attach to the frame is 60° for a three leaflet valve.
U.S. Pat. No. 5,500,016 (Fisher) entitled “Artificial Heart Valve” discloses a valve having a leaflet shape defined by the mathematical equation z2+y2=2RL (x−g)−α(x−g)2, where g is the offset of the leaflet from the frame, RL is the radius of curvature of the leaflet at (g,0,0) and α is the shape parameter and is >0 and <1.
A valve design having a partially open configuration when the valve is not subject to a pressure gradient, but assuming a fully-open position during forward flow is disclosed in International Patent Application WO 97/41808 entitled “Method for Producing Heart Valves”. The valve may be a polyurethane trileaflet valve and is contained within a cylindrical outer sleeve.
U.S. Pat. Nos. 4,222,126 (Boretos et al.) and 4,265,694 (Boretos et al.) disclose a trileaflet polyurethane valve with integral polyurethane elastomeric leaflets having their leading edges reinforced with an integral band of polymer and the leaflets reinforced radially with thicker lines of polyurethane.
The problem of chronic thrombus formation and tissue overgrowth arising from the suture ring of valves has been addressed by extension of the valve body on either side of the suture ring as disclosed in U.S. Pat. No. 4,888,009 (Lederman et al.) entitled “Prosthetic Heart Valve”.
Current polyurethane valve designs have a number of potential drawbacks. Close coaptation of leaflets, while ensuring good valve closure, limits the wash-out of blood during hemodynamic function, particularly in the regions close to the stent posts at the commissures. This region of stagnation is likely to encourage local thrombogenesis, with further restriction of the valve orifice in the longer term as well as increasing the risk of material embolising into the circulation. Associated with the thrombosis may be material degradation (in non-biostable polyurethanes) and calcification resulting in localised stiffening the leaflets, stress concentrations and leaflet failure. As previously discussed, animal implants of a trileaflet polyurethane valve design have indicated that thrombus does tend to collect in this region, restricting the valve orifice and damaging the structure of the valve.
Present valve designs are limited by the availability of suitable polyurethanes which possess good mechanical properties as well as sufficient durability to anticipate clinical functionality of up to twenty years or more. Many low modulus materials, which provide good hydrodynamic function, fail during fatigue testing at unacceptably low durations, due to their greater susceptibility to the effects of accumulated strain. Higher modulus polyurethanes may be better able to withstand repeated stress without accumulating significant damage, but are too stiff to provide good hydrodynamic function in conventional almost-closed geometry valve designs. Current design strategies have not been directed towards enabling the incorporation of potentially more durable, higher modulus leaflet materials, nor the creation of a valve design that is able to maintain good hydrodynamic function with low modulus polyurethanes manufactured as thick leaflets.
The nature of the valve leaflet attachment to the frame is such that, in many valve designs, there is a region of leaflet close to the frame, which is restrained by the frame. This region may extend some distance into the leaflet before it interfaces with the free-moving part of the leaflet, or may be directly at the interface between frame and leaflet. There thus exists a stress concentration between the area of leaflet that is relatively mobile, undergoing transition between fully open and fully closed, and the relatively stationary commissural region. The magnitude of this flexural stress concentration is maximized when the design parameters predicate high bending strains in order for the leaflet to achieve its fully open position.
U.S. Pat. Nos. 4,222,126 (Boretos et al.) and 4,265,694 (Boretos et al.) disclose a valve which uses thickened leaflet areas to strengthen vulnerable area of the leaflets. However this approach is likely to increase the flexure stress and be disadvantageous in terms of leaflet hydrodynamic function.
The major difficulties which arise in designing synthetic leaflet heart valves can be explained as follows. The materials from which the natural trileaflet heart valves (aortic and pulmonary) are formed have deformation characteristics particularly suited to the function of such a valve. Specifically, they have a very low initial modulus, and so they are very flexible in bending, which occurs at low strain. This low modulus also allows the leaflet to deform when the valve is closed and loaded in such a way that the stresses generated at the attachment of the leaflets, the commissures, are reduced. The leaflet material then stiffens substantially, and this allows the valve to sustain the closed loads without prolapse. Synthetic materials with these mechanical properties are not available.
Polyurethanes can be synthesized with good blood handling and good durability. They are available with a wide range of mechanical properties, although none has as low a modulus as the natural heart valve material. Although they show an increase in modulus at higher strains, this does not occur until strains much higher than those encountered in leaflet heart valves.
Polyurethanes have been the materials of choice for synthetic leaflet heart valves in the last decade or more. More recently, polyurethanes have become available which are resistant to degradation when implanted. They are clearly more suitable for making synthetic leaflet heart valves than non-stable polyurethanes, but their use suffers from the same limitations resulting from their mechanical properties. Therefore, design changes must be sought which enable synthetic trileaflet heart valves to function with the best available materials.
Key performance parameters which must be considered when designing a synthetic leaflet heart valve include pressure gradient, regurgitation, blood handling, and durability.
To minimize the gradient across the open valve, the leaflets must open wide to the maximum orifice possible, which is defined by the inside diameter of the stent. This means that there must be adequate material in the leaflets so they can be flexed into a tube of diameter equal to the stent internal diameter. In addition, there has to be a low energy path for this bending because the pressure forces available to open the valve are small, and the lower the gradient, the smaller the pressure becomes. All the leaflets must open for the lowest cardiac output likely to be encountered by that valve in clinical service.
To minimize closing regurgitation (reverse flow lost through the closing valve) the valve leaflets must be produced at or close to the closed position of the valve. To minimize closed valve regurgitation (reverse flow through the valve once it has closed), the apposition of the leaflets in the commissural region is found to be key, and from this perspective the commissures should be formed in the closed position.
Proper blood handling means minimising the activation both of the coagulation system and of platelets. The material of construction of the valve is clearly a very important factor, but flow through the valve must also avoid exposing blood either to regions of high shear (velocity gradient) or to regions of relative stasis. Avoiding regions of high shear is achieved if the valve opens fully, and relative stasis is avoided if the leaflet/frame attachment and the commissural region in particular opens wide. This is not achieved with typical synthetic materials when the commissures are molded almost closed, because the stiffness of synthetics is too high.
Durability depends to a large extent on the material of construction of the valve leaflets, but for any given material, lifetime will be maximized if regions of high stress are avoided. The loads on the closed valve are significantly greater than loads generated during valve opening. Therefore, the focus should be on the closed position. Stresses are highest in the region of the commissures where loads are transmitted to the stent, but they are reduced when the belly of the leaflet is as low as practicable in the closed valve. This means that there must be sufficient material in the leaflet to allow the desired low closing.